The present disclosure relates to systems and methods for emission tomography and, more particularly, to systems and methods for emission tomography that provide an improvement in image quality when using advanced radionuclides, wherein such radionuclides are characterized by the emission of both prompt gamma rays and positrons upon decay.
There are a variety of emission tomography imaging systems and methods. One clinically important example is positron emission tomography (PET) which, generally, utilizes an administered radionuclide (also considered a radioactive isotope or radioisotope) to acquire two-dimensional and three-dimensional tomographic images of a target area or organ of interest in a subject. More specifically, such radionuclides are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances, such as glucose or carbon dioxide. These radiopharmaceuticals are then administered to the patient where they become involved in biological processes such as blood flow; fatty acid and glucose metabolism; and protein synthesis. Through a respective biological process, the radiopharmaceuticals accumulate in, or otherwise target, the area or organ of interest in the subject. By measuring or identifying photons emitted from the area or organ of interest by the accumulated or targeted radiopharmaceutical, clinically useful biological and physiological information can be acquired.
Conventional radionuclides such as fluorine-18 (18F), carbon-11 (11C), nitrogen-13 (13N), and oxygen-15 (15O) are commonly used to label PET radiopharmaceuticals. These radionuclides are denominated “conventional” because they decay by emitting only positrons. The positrons travel a very short distance before they encounter an electron and, when this occurs, the positrons are annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features that are pertinent to PET imaging. Namely, each annihilation gamma ray has an energy of 511 keV and the two gamma rays are directed in substantially opposite directions. An image is created by determining the number of such annihilation events at each location within the scanner's field of view.
To create such an image, typical PET scanners consist of one or more rings of detectors which are positioned to encircle the subject. Coincidence detection circuits connected to the detectors record only those photons that are detected simultaneously by two detectors located on opposite sides of the subject and that fall within an energy acceptance window around 511 keV. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes, millions of events can be recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well-known tomographic reconstruction techniques.
More specifically, current clinical (and most preclinical) PET scanners and systems include a ring 100 of block detectors 102 for detecting emitted photons, typically in circular, such as the array shown in FIG. 1, or in hexagonal or octagonal arrays. Block detectors 102 include a piece of scintillator material that converts the energy deposited by gamma rays into visible light. The scintillator material is usually segmented into many scintillation crystal elements configured in an array, which is read out by one or more photon detectors (such as a number of individual photo-multiplier tubes (PMTs), a position-sensitive photo-multiplier tube (PS-PMT), or silicon photo-multipliers (Si-PM)) that convert the light emitted by the scintillation material into electrical signals whose magnitude is proportional to the energy deposited by the gamma rays in the scintillator material. By combining the output signal of the photon detector(s) of the block detector, it is possible to determine the single crystal in which the detected photon interacted and the energy deposited by such photon.
Furthermore, as shown in FIG. 1, the ring 100 of block detectors of a PET scanner includes individual detectors that are operated in coincidence with a fan beam 104 of block detectors on the opposite side of the ring 100. The inner circle 106 formed by edges of all such fan beams defines the useful field of view. Data is usually recorded simultaneously for all possible fan beams, and the PET scanner will produce an output whenever two photons are detected in opposite block detectors of a fan beam 104 within a specified coincidence timing window (for example, in the range of hundreds of picoseconds to tens of nanoseconds) and when both events fall into a predetermined energy window (usually from 511 keV−ΔE1 to 511 keV+ΔE2, where ΔE1 and ΔE2 are a function of the energy resolution of the block detectors). Any such events are called double coincidences.
In addition to the conventional radionuclides described above, advanced PET radionuclides (also considered non-standard radionuclides), for example, such as iodine-124 (124I), bromine-76 (76Br), yttrium-86 (86Y), among others, may be useful for preclinical and clinical studies due to their chemical properties and their relatively long half-life. These properties make them especially well-suited for labeling antibodies, for dosimetry in internal radiotherapy procedures, and for an easy distribution from a cyclotron where they are generated to distant imaging centers. Other advanced radionuclides like rubidium-82 (82Rb), which is currently used in cardiac studies, have a short half-life, but can be obtained from a generator, as a decay product of a long half-life parent radionuclide.
Most of these advanced radionuclides, however, have a drawback because, when they decay, prompt gamma rays can be emitted in addition to positrons. This causes the emission of more than two gamma rays per radioactive decay and, as the energy of the additional prompt gamma rays or scattered photons from them may be close to the energy of the standard annihilation gamma rays (that is, about 511 keV), it is difficult to distinguish the prompt gamma rays from the annihilation gamma rays (as the energy resolution of existing PET scanners ranges from about 10% to about 30%). Also, in some cases, the prompt gamma ray has an energy significantly larger than 511 keV, but it may deposit in the detector only part of its energy. If the measured energy is close to about 511 keV, the detected prompt gamma ray cannot be distinguished from the standard gamma rays. As a result, spurious double coincidences may be detected in a conventional PET scanner, causing a significant additional background in reconstructed images, reducing image contrast, decreasing the detectability of hot spots in the images, and compromising their quantitative properties. Such coincidences may be referred to as non-standard spurious coincidences or simply spurious coincidences, for example as opposed to standard coincidences further described below. These coincidences are considered spurious in a spatial or geometrical context, as opposed to a temporal context, because their resulting line of response, described below, does not pass through the point of annihilation.
More specifically, commercial PET scanners are designed to detect and record only double coincidences (as opposed to triple and/or other multiple coincidences that involve the detection, in coincidence, of more than two gamma-rays). The data from these double coincidences are usually stored in a large list of events or in a histogram format (such as a sinogram or line-of-response histograms). When such advanced radionuclides (considered positron plus prompt gamma ray emitters) are used, a number of different events may be detected by the PET scanner. For example, as shown in FIG. 2A, the two annihilation gamma rays A′, B′ can be detected along their correct line of response (that is, line A-B). This is considered a standard coincidence and, more specifically, a true coincidence (other types of standard coincidences can include random coincidences or in-body scatter coincidences). However, as shown in FIG. 2B, a prompt gamma ray C′ can be detected in coincidence with one of the annihilation gamma rays A′, resulting in an incorrect, or spurious, line of response (that is, line A-C). This may occur, for example, when gamma ray B′ escapes from the scanner gantry or only deposits a portion of its energy at the detector due to scattering, as shown in FIG. 2B, resulting in the scanner not detecting gamma ray B′ in coincidence with gamma ray A′. This is considered a non-standard spurious coincidence. Thus, in advanced radionuclide applications, when detected double coincidences are reconstructed using standard reconstruction methods, a significant background in the image can be noticed due to the non-standard spurious coincidences between one annihilation gamma ray and one prompt gamma ray, like that shown in FIG. 2B.
Several methods have been proposed for correcting the spurious activity by estimating the background caused by prompt gamma rays and removing this background during image reconstruction. These methods involve, for example, subtraction of a uniform distribution with an intensity that is obtained using the outer region of the field-of-view (where no real activity is expected to be present) as a reference, modification of the parameters of traditional scatter correction processes, and/or a convolution subtraction method (similar to scatter correction) using either empirically or analytically determined kernel functions. These methods either combine scatter correction and spurious-coincidence-background correction into a single correction assuming a roughly fixed relationship between the two or they try to model the background distribution. Nevertheless, such approaches are often not accurate because the shape of the spurious coincidence distribution can be very different from that of scatter and their relative magnitudes vary with the size, shape, and density of the object being imaged and the radionuclide used.
For example, a first proposed method includes estimating the background created by spurious coincidences by projecting in random directions from points inside measured lines of response. This method obtains an estimation that is later used for subtracting the background from signal data. The limitations of this method are that it requires significant additional computational time for a realistic estimation and that it can fail if the model for the projection is not accurate and not representative of the actual scanner.
In a second proposed method, the first method is combined with a scatter estimation to obtain an additional estimation of the contribution of the spurious background. This method requires some measurement in regions where it is assumed there should be no signal in order to scale the estimation. However, this is not possible for some cases, such as when scanning obese patients that fill the whole or nearly the whole field-of-view of the scanner.
In general, these estimation methods are time-consuming, can introduce bias into the images, depend on the size of the subject, and can affect the statistical properties of the reconstructed images. As a result, they are not effective solutions for improving image quality in PET when using advanced radionuclides and, furthermore, can often have a negative effect on resulting images.
Therefore, it would be desirable to have a system and method to provide a direct measurement of the background caused by prompt gamma rays during PET imaging with advanced radionuclides in order to improve the image quality without relying on inaccurate, time-consuming estimations based on simulations or approximations.